It is often desirable to image biological tissue through intervening tissues or structure, for example, through overlying light transmissive layers of cells (e.g., in the breast) or through fluids (e.g., the aqueous or vitreous humor in the eye). Imaging through intervening tissue or structure allows tissue to be studied in relatively thick sections or in vivo.
To a limited extent, such imaging of internal structures may be done using a conventional microscope by focusing the microscope objective “through” the overlying layers so that the structure of interest is at the focal plane of the microscope objective and sharply in focus and other overlying structures are defocused.
Confocal microscopy takes this process a step further by placing a light stop in the optical path to block all light not received from the single focal spot of the microscope objective. Scanning the focal spot through the tissue and measuring variations of brightness as a function of that scan, can produce an image free from light interference from adjacent layers in the tissue. Unfortunately, the optical stop significantly limits the light through the confocal microscope, requiring a bright light source usually provided by a laser and long exposure times.
Recently developed techniques allow virtually any protein in a cell to be tagged with fluorescent molecules. The fluorescent molecules, and thus the tagged cells, can then be visualized by exciting the fluorescent molecule with an excitation light beam. The excitation beam is typically of a different frequency than the frequency of fluorescence so that a dichroic filter can be used to block the excitation beam, making the tagged tissue stand out.
Referring to FIG. 1, an improved variation on confocal microscopy makes use of this fluorescent tagging in a process called multi-photon fluorescence. In multi-photon fluorescence, a fluorescent molecule 10 may simultaneously absorb two (or more) photons 12 to move to an excited state 14 elevated by at least twice the energy of each individual photon 12. A subsequently emitted fluorescence 16 will have approximately twice the frequency of the stimulating photons 12 to be readily distinguishable from the photons 12 of the exciting beam. Importantly, the property of multi-photon fluorescence is nonlinearly related to light intensity and thus multi-photon fluorescence can be controlled to occur in only small regions where the excitation light beam is focused to an intensity causing significant multi-photon fluorescence. Tissue before and after this focused region, even if tagged by the fluorescent molecules, will exhibit only weak multi-photon fluorescence.
Referring to FIG. 2, a multi-photon microscope 20, exploiting this principal, typically employs a light source 22 and provides an excitation beam 23 of stimulating photons 12 which are then received by an optical assembly 24 which focuses the beam 23 at a focal plane 26 to a focal spot 30. As the beam 12 narrows with focusing, the intensity increases and the amount of multi-photon fluorescence 32 increases rapidly causing the tissue to fluoresce principally only at the focal spot 30 in the focal plane 26. Light 35 from that fluorescence passes backward through the optical assembly 24 and is reflected off a dichroic mirror 36 separating it from an excitation beam 23 to be received by a photodetector 38. The spot 30 is scanned through tissue in a three-dimensional raster pattern 40, and brightness values obtained by the photodetector 38 are mapped to the locations in the tissue to provide the ability to reconstruct images of embedded structures in the tissue free from the influence of underlying or overlying tissue.
Such multi-photon fluorescence techniques have been used to provide sharp images of in vivo tissue up to a depth of about 600 μm. Beyond this depth, the ability to provide a small focal spot 30 (which ultimately determines the resolution of the image) degrades because of inhomogeneities in the optical properties of the intervening tissue, principally refractive index, which distort the incident waveform preventing sharp focus.
The principles of adaptive optics have been applied to correct the problem of wavefront distortion. Here the goal is to pre-distort the wavefront of the excitation beam to exactly offset the aberration caused by the intervening tissue. Such approaches may use deformable mirrors which have a continuous surface electrically flexed to change local elevation of the surface and thereby advance or retard a wavefront reflected from that surface, by precise amounts. Alternative approaches use liquid crystal devices (LCDs) which change an index of refraction as a function of voltage over their surface, for example, by using LCDs as Fresnel lenses.
Such LCD devices are relatively slow with low contrast and power handling capabilities while deformable mirrors are extremely costly and/or of relatively low resolution. The amount of phase shift achievable in a deformable mirror is severely limited by the small deformation range and the constraints imposed by a continuous mirror surface. Limitations in phase shift range prevent such devices from producing the significant phase shifts necessary to accommodate phase distortions incident to imaging structure deep within tissue. For the deformable mirror, the deflection range is smaller for higher resolution devices, effecting an undesired trade-off between the imaging depth and resolution.